The present disclosure relates to a method for irradiating an irradiation target with a non-continuous particle beam, in particular a proton beam. Furthermore, the present disclosure relates to an apparatus for carrying out said method.
Charged particle beams consisting of protons or heavier ions are successfully used in cancer therapy to destroy tumors by irradiation. One of the advantages of such charged particle beams is the Bragg peak at the end of the beam path where a large fraction of the irradiation dose is deposited. The depth of the Bragg peak can be varied by varying the particle beam energy. This allows for control of the irradiation depth and, in combination with lateral beam spreading or deflection, a good three dimensional dose conformation, i.e. an effective delivery of the dose to the target volume (tumor) while avoiding damages in neighboring regions (healthy tissue).
A charged particle therapy system is for example described in U.S. Pat. No. 5,260,581. Charged particle therapy systems usually comprise an accelerator for producing a charged particle beam, a beam transport system for transporting the beam to the patient, and various means for treating the beam in order to achieve a good conformation of the dose to the target volume. The term “treating the beam” shall encompass all possible forms of manipulating the beam, including, without limitation, focusing, spreading, deflecting and interrupting the beam and changing the energy of the beam particles.
Various types of particle accelerators are used in charged particle therapy, inter alia synchrotrons and cyclotrons. In synchrotrons, particles are accelerated in an orbit resulting from a magnetic field that is actively changed with time to keep the orbital radius constant. The beam extracted from a synchrotron is not continuous but pulsed in discrete bunches. The discrete beam bunches are extracted from a synchrotron with the same frequency as with which the magnetic field is cycled. By a method called “slow extraction”, the beam bunch accelerated in a synchrotron can be gradually extracted over time. Elongated pulses can be extracted such that a semi-continuous beam can be generated. During the time of slow extraction, which can be several seconds long, the extracted beam is continuous, but after extraction of the accelerated bunch, the magnetic field must be cycled again and the beam is interrupted. A synchrotron allows the energy of the extracted particles to vary over a broad range.
A cyclotron is a circular accelerator which has a magnetic field constant over time but with the magnetic field strength changing with the radius. A particle accelerated in a cyclotron moves on a spiral path with increasing radius in the plane normal to the magnetic field. Particles are accelerated in a cyclotron by applying an alternating radio frequency (RF) voltage to one or more electrodes, called “dees”. The RF voltage generates an electric filed across the gap between the adjacent dees. The orbital period of the charged particles in the magnetic field must be synchronized to the RF voltage so that the particles are effectively accelerated as they repeatedly cross the dee gaps. The synchronization must be adjusted in such a way that also relativistic mass effects are compensated for.
Two different types of cyclotrons have been developed that solve the synchronization between orbital period and RF voltage in different ways. The “isochronous cyclotron” uses a constant frequency of the voltage and has a magnetic field that increases with the radius. The shape of the magnetic field compensates for the relativistic mass increase of the charged particles with acceleration. Thereby, the isochronous cyclotron is capable of producing a continuous beam without interruptions.
In a synchrocyclotron the magnetic field is constant or decreasing with increasing radius of the accelerator and the RF frequency of the accelerating voltage is adjusted to achieve synchronization with the orbital period of the charged particles. The RF frequency of the acceleration is changed (modulated) in a cycle, starting at the highest or “injection” frequency and decreasing over time to the lowest frequency or “extraction” frequency. After reaching the extraction frequency, the modulation cycle is started again with the injection frequency. As a consequence, a synchrocyclotron can only accelerate one discrete bunch of charged particles per RF frequency modulation cycle to the final accelerator energy. The time structure of a particle beam extracted from a synchrocyclotron is pulsed as only particles are extracted when the RF frequency of the accelerating voltage is equal to the extraction frequency. The length of the particle bunch from a synchrocyclotron is typically in the order of 105 times shorter than a bunch extracted from a synchrotron using slow extraction.
In all types of accelerators described above, the RF accelerating voltage imposes a “micro-structure” on the beam in the longitudinal direction. The inverse of the frequency or period of the accelerating voltage is orders of magnitude smaller than the time constants relevant for operating charged particle therapy apparatus, such as the setting time of scanning magnets and energy degraders, the measurement time needed by beam diagnostic means, or the bunch length produced by synchrotrons and synchrocyclotrons or other accelerators. Therefore, the RF micro-structure of the beam or beam-bunches is neglected in the remainder of this document.
In charged particle therapy various methods are known to achieve a good dose conformation to the target volume. They are usually grouped into “passive” methods and “active” methods which are inter alia described in W. Chu et al. (Rev. Sci. Instrum. 64, pp. 2055 (1993)). Passive systems generally use scattering systems in order to broaden the beam and to cover the treatment areas required in charged particle therapy. However, expensive patient specific equipment is required and neutrons generated in the scattering systems lead to unwanted increased neutron doses for the patient.
According to active methods, for example described in W. Chu et al. (Rev. Sci. Instrum. 64, pp. 2055 (1993)), the beam is deflected and scanned over the target, for example by use of deflection magnets. Some active methods require patient specific path compensators, bolus or collimators which increase effort and cost of the treatment. This is avoided by an active method known as “pencil beam scanning” or “spot scanning” which is described in E. Pedroni et al. (Med. Phys. 22 (1), 1995). Pencil beam scanning uses continuous or semi-continuous beams. The irradiation spot is moved in the distal direction, i.e. in the direction of the beam, by changing the energy of the beam, most commonly by using an energy degrader. The movement of the beam in the X- and Y-direction in the plane normal to the direction of the beam is performed with two scanning magnets. By using a focused “pencil beam” individual small volumes, also referred to as “voxels”, can be treated. The whole tumor is subdivided into voxels and then irradiated voxel by voxel. During treatment, the beam is moved to a specific voxel, and this voxel is irradiated until a dose monitor detects that the required dose level for this voxel has been reached. The (semi-)continuous beam is turned off, and the machine parameters are adjusted for the next voxel. In order to avoid imprecise dose application, the beam switch has to be fast and exact and the intensity has to be kept at a lower level which leads to lengthy treatment times. Several techniques are used in pencil beam scanning to tune the motion and the intensity of the beam in order to regulate the dose applied to each voxel. In some methods, the movement of the beam in three dimensions is performed without interrupting the beam. In other methods, the beam is interrupted after the irradiation of a single voxel. In this case, the irradiation is re-started after the two scanning magnets and the energy degrader have reached the settings needed for the next voxel. However, in all these methods a continuous or semi-continuous beam, for example from an isochronous cyclotron or a synchrotron with slow extraction, is used to irradiate a single voxel.
The use of a non-continuous (pulsed) beam with known pencil beam techniques leads either to very long treatment times or to imprecise dose application. The bunch length of a non-continuous beam is usually too short to be interrupted or subdivided. Therefore, if the required dose precision per voxel is +/−2%, a single bunch of particles can only have 2% of the total dose per voxel, so that 50 bunches are needed per voxel. This leads to undesirable long treatment times. Furthermore, the techniques used to avoid imprecisions due to organ motion lengthen treatment times. One of such techniques is gating, i.e. setting a treatment window such that irradiation is only applied during certain phases of organ motion, e.g. during certain phases of the breathing cycle of the patient. Another technique to mitigate the effects of organ motion is to average out the movements by irradiating all voxels multiple times. All of these techniques lead to even longer treatment times.
Pencil beam scanning has several advantages with respect to other active or passive methods. The dose conformation can be significantly improved, the neutron dose for the patient can be decreased and the time consuming and costly process of production and mounting of patient specific compensators and collimators can be omitted. However, according to current techniques, pencil beam scanning can only be performed with continuous or semi-continuous beams. The particle accelerators typically needed to produce such (semi-)continuous beams, for example isochronous cyclotrons and synchrotrons with slow extraction, are usually mechanically larger than accelerators producing a non-continuous or pulsed particle beam such as synchrocyclotrons or linear accelerators (linacs).